Conventional ultrasound imaging systems typically include a hand-held scan head coupled by a cable to a large rack-mounted console processing and display unit. The scan head typically includes an array of ultrasonic transducers which transmit ultrasonic energy into a region being imaged and receive reflected ultrasonic energy returning from the region. The transducers convert the received ultrasonic energy into low-level electrical signals which are transferred over the cable to the processing unit. The processing unit applies appropriate beam forming techniques such as dynamic focusing to combine the signals from the transducers to generate an image of the region of interest.
Typical conventional ultrasound systems can have transducer arrays which consist of 128 ultrasonic transducers. Each of the transducers is associated with its own processing circuitry located in the console processing unit. The processing circuitry typically includes driver circuits which, in the transmit mode, send precisely timed drive pulses to the transducer to initiate transmission of the ultrasonic signal. These transmit timing pulses are forwarded from the console processing unit along the cable to the scan head. In the receive mode, beam forming circuits of the processing circuitry introduce the appropriate delay into each low-level electrical signal from the transducers to dynamically focus the signals such that an accurate image can subsequently be generated.
A schematic block diagram of an imaging array 18 of N piezoelectric ultrasonic transducers 18(1)-18(N) as used in an ultrasound imaging system is shown in FIG. 1A. The array of piezoelectric transducer elements 18(1)-18(N) generate acoustic pulses which propagate into the image target (typically a region of human tissue) or transmitting media with a narrow beam. The pulses propagate as a spherical wave with a constant velocity. Acoustic echoes in the form of returning signals from image points P or reflectors are detected by the same array 18 of transducer elements or another receiving array and can be displayed in a fashion to indicate the location of the reflecting structure P.
The acoustic echo from the point P in the transmitting media reaches each transducer element 18(1)-18(N) of the receiving array after various propagation times. The propagation time for each transducer element is different and depends on the distance between each transducer element and the point P. This holds true for typical ultrasound transmitting media, i.e. soft bodily tissue, where the velocity of sound is assumed (or relatively) constant. Thereafter, the received information is displayed in a manner to indicate the location of the reflecting structure.
In two-dimensional B-mode scanning, the pulses can be transmitted along a number of lines-of-sight as shown in FIG. 1A. If the echoes are sampled and their amplitudes are coded as brightness, a grey scale image can be displayed on a CRT. An image typically contains 128 such scanned lines at 0.75.degree. angular spacing, forming a 90.degree. sector image. Since the velocity of sound in water is 1.54.times.10.sup.5 cm/sec, the round-trip time to a depth of 16 cm will be 208 .mu.s. Thus, the total time required to acquire data along 128 lines of sight (for one image) is 26.6 ms. If other signal processors in the system are fast enough to keep up with this data acquisition rate, two-dimensional images can be produced at rates corresponding to standard television video. For example, if the ultrasound imager is used to view reflected or back scattered sound waves through the chest wall between a pair of ribs, the heart pumping can be imaged in real time.
The ultrasonic transmitter is typically a linear array of piezoelectric transducers 18(1)-18(N) (typically spaced half-wavelength apart) for steered arrays whose elevation pattern is fixed and whose azimuth pattern is controlled primarily by delay steering. The radiating (azimuth) beam pattern of a conventional array is controlled primarily by applying delayed transmitting pulses to each transducer element 18(1)-18(N) in such a manner that the energy from all the transmitters summed together at the image point P produce a desired beam shape. Therefore, a time delay circuit is needed in association with each transducer element 18(1)-18(N) for producing the desired transmitted radiation pattern along the predetermined direction.
For a given azimuth angle, as can be seen in FIG. 1B, there can be two different transmitting patterns: a "single-focus" and a "zone-focus" pattern. The single-focus method employs a single pulse focused at mid-range of the image line along a particular line of sight. In a single pulse mode, the azimuth focus depth can be electronically varied, but remains constant for any predetermined direction. In zone-focus operation, multiple pulses, each focused at a different depth (zone), are transmitted along each line of sight or direction. For multiple pulse operation, the array of transmitters is focused at M focal zones along each scan direction, i.e., a series of M pulses is generated, P.sub.0, P.sub.1, . . . , P.sub.M-1, each pulse being focused at its corresponding range R.sub.0, . . . , R.sub.M-1, respectively. The pulses are generated in a repeated sequence so that, after start up, every Mth pulse either begins a look down a new direction or corresponds to the initial pulse P.sub.0 to repeat the series of looks down the present direction. For the zone-focused mode, a programmable time-delay circuit is needed in association with each transducer element to produce beam patterns focused at different focal zones.
As previously described, the same array 18 of transducer elements 18(1)-18(N) can be used for receiving the return signals. The reflected or echoed beam energy waveform originating at the image point reaches each transducer element after a time delay equal to the distance from the image point to the transducer element divided by the assumed constant speed of the waveform of signals in the media. Similar to the transmitting mode, this time delay is different for each transducer element. At each receiving transducer element, these differences in path length should be compensated for by focusing the reflected energy at each receiver from the particular image point for any given depth. The delay at each receiving element is a function of the distance measured from the element to the center of the array and the viewing angular direction measured normal to the array. It should be noted that in ultrasound, acoustic pulses generated by each transducer are not wideband signals and should be represented in terms of both magnitude and phase.
The beam forming and focusing operations involve forming a sum of the scattered waveforms as observed by all the transducers, but in this sum, the waveforms must be differentially delayed so that they will all arrive in phase and in amplitude in the summation. Hence, a beam forming circuit is required which can apply a different delay on each channel, and vary that delay with time. Along a given direction, as echoes return from deeper tissue, the receiving array varies its focus continually with depth. This process is known as dynamic focusing.
FIGS. 2A-2C show schematic block diagrams of three different conventional imaging or beam focusing techniques. A non-programmable physical lens acoustic system 50 using an acoustic lens 51 is shown in FIG. 2A. In turn, dynamic focusing systems where associated signal processing electronics are employed to perform real-time time delay and phase delay focusing functions are respectively shown in FIGS. 2B and 2C. FIG. 2B shows a time delay system 52 using time delay elements 53, and FIG. 2C shows a phase delay system 54 using phase delay elements 55.
In the lensless systems of FIGS. 2B and 2C, the signal processing elements 53, 55 are needed in association with each receiving transducer element, thus defining processing channels, to provide time delay and focus incident energy from a field point to form an image. Accordingly, a beam forming circuit is required which can provide a different delay on each processing channel, and to further vary that delay with time. Along a predetermined direction, as echoes return from distances further away from the array of transducer elements, the receiving array varies its focus continually with depth to perform dynamic focusing.
After the received beam is formed, it is digitized in a conventional manner. The digital representation of each received pulse is a time sequence corresponding to a scattering cross section of ultrasonic energy returning from a field point as a function of range at the azimuth formed by the beam. Successive pulses are pointed in different directions, covering a field of view from -45.degree. to +45.degree.. In some systems, time averaging of data from successive observations of the same point (referred to as persistence weighting) is used to improve image quality.
For example, in an ultrasound imaging system operating at a 2-5 MHz frequency range, an electronic circuit capable of providing up to 10 to 20 .mu.s delay with sub-microsecond time resolution is needed for the desired exact path compensation. As shown in FIG. 2B, a delay line is inherently matched to the time-delay function needed for dynamic focusing in a lensless ultrasound system.
More specifically, in an exemplary ultrasound imaging system with a 5 MHz operating frequency and an array of 128 transducer elements on half-wavelength centers, a straightforward delay implementation requires each processing channel/transducer element to include either a 480-stage delay line with a clock period programmable with a 25 ns resolution or a 480-stage tapped delay line clocked at 40 MHz in conjunction with a programmable 480-to-one time-select switch to set the appropriate delay. There are two problems associated with these conventional techniques. First, a simple variable-speed clock generator has not been developed to date. Secondly, for an N-stage tapped delay line, the area associated with the tap select circuit is proportional to N.sup.2, thus such a circuit would require a large amount of microchip area to realize an integrated tap architecture.
Due to the difficulty and complexity associated with the generation of the control circuits of the conventional approach, only a few time-delay structures could be integrated on one microchip, and therefore a large number of chips would be needed to perform a multi-element dynamic beam forming function. For these reasons, none of the prior art ultrasound imaging systems utilize the straightforward time-delay implementation. Instead, a plane-wave mixer approximation is used. In this approximation process, the total delay is separated into two parts: an analog plane-wave mixer technique is used to approximate the required fine delay time and a true coarsely spaced delay line is used to achieve the coarse delay time.
In accordance with the plane-wave approximation, the fine delay can be achieved by modifying the phase of AC waves received by each receiving processing channel and implemented by heterodyning the received waves from each receiving transducer element with different phases of a local oscillator, i.e., creating analog phase shifting at each processing channel. Specifically, by selecting a local oscillator with a proper phase angle of the form cos (.omega..sub.o t+.OMEGA..sub.n (t)), where .OMEGA..sub.n is chosen to satisfy the expression .OMEGA..sub.n (t)=.omega..sub.o (T'.sub.n (t)-T'.sub.n (t)), T.sub.n (t) is the ideal compensating delay and T'.sub.n (t) is a coarsely quantized approximation of T.sub.n. It will be appreciated that when the mixer output is delayed by T'.sub.n, the phase of one of its intermediate frequency (IF) sidebands provides phase coherence among all the processing channels.
In the conventional implementation of the aforementioned technique, a tap select is used which connects any received down-conversion mixer output to any tap on a coarsely spaced, serially connected delay line. The tap select is essentially a multiposition switch that connects its input to one of a number of output leads. One output lead is provided for each tap on the delay line. Therefore, each mixer output can be connected to a few coarsely spaced taps on a delay line, and all the tap outputs can be summed together coherently. However, for an exemplary 5 MHz operation, if a single-mixer arrangement as described above is used, a delay line with delay resolution less than one microsecond is needed.
In summary, the conventional technique described heretofore involves heterodyning the received signals with an oscillator output by selecting a local oscillator frequency so as to down convert the output to an IF frequency. This down converted signal is then applied to another mixer. By selecting the proper phase angle of the second oscillator, the phase of the intermediate frequency waves produced by the second heterodyning is controlled. The output of the second mixer is then connected through a tap select to only one, or at most a few, coarsely spaced taps on a delay line during the focal scanning along each direction.
The aforementioned approximation technique is used due to the fact that given an image that is somewhat out of focus, the image can be focused in an economically feasible manner by utilizing readily available techniques such as analog mixers and RC networks. Unfortunately, the mixer approximation method suffers from image misregistration errors as well as signal loss relative to the ideally-focused (perfect delay) case.
Modern ultrasound systems require extensive complex signal processing circuitry in order to function. For example, hundreds of delay-and-sum circuits are needed for dynamic beam forming. Also, pulsed or continuous Doppler processors are needed for providing two-dimensional depth and Doppler information in color flow images, and adaptive filters are needed for clutter cancellation. Each of these applications requires more than 10,000 MOPS (million operations per second) to be implemented. Even state-of-the-art CMOS chips only offer several hundred MOPS per chip, and each chip requires a few watts of electric power. Thus, an ultrasound machine with a conventional implementation requires hundreds of chips and dissipates hundreds of watts of power. As a result, conventional systems are implemented in the standard large rack-mounted cabinets.
Another drawback in conventional ultrasound systems is that the cable connecting the scan head to the processing and display unit is required to be extremely sophisticated and, hence, expensive. Since all the beam forming circuitry is located in the console, all of the low-level electrical signals from the ultrasonic transducers must be coupled from the scan head to the processing circuitry. Because the signals are of such a low level, they are extremely susceptible to noise, crosstalk and loss. With a typical transducer array of 128 transducers, the cable between the scan head and the processing and display console is required to contain 128 low-noise, low-crosstalk and low-loss coaxial cables. Such a cable requires expensive materials and extensive assembly time and is therefore very expensive.